Diagnostic ultrasound is an essential modality in virtually every medical specialty and particularly in obstetrics, cardiology and radiology. The ultrasound transducer is the critical component and the limiting factor affecting the quality of diagnostic ultrasound imaging and Doppler measurements. In a conventional circular piston piezoelectric transducer used in mechanical scanning for medical applications (e.g., 19 mm diameter, 3.5 MHz resonant frequency) the electrical impedance of the transducer is approximately 50.OMEGA.. Such a transducer is well matched to the conventional electrical transmit circuit for delivering large amounts of acoustic power to the tissue load during the transmit mode. In a like manner, in receive mode, such a transducer is well suited for driving the typical 50.OMEGA. or 75.OMEGA. coaxial cable connected to the amplifier circuits of the scanning system.
The most sophisticated medical ultrasound scanners now typically use (N.times.1) linear arrays containing over a hundred transducer elements which may be multiplexed and/or electronically steered and focused via phased array techniques. A linear array transducer for phased array scanning, typically employs much smaller array elements than the conventional transducer described above. For example, a typical linear array includes 128 elements, each of which is 0.2 mm wide by 10 mm long with a resonant frequency of 3.5 MHz. Each piezoelectric ceramic transducer element in a linear array acts as a capacitor of approximately 150 picofarads (pf), which produces an electrical impedance Z=R.sub.a +jX where X.apprxeq.300.OMEGA. at 3.5 MHz. At resonance the reactive component is in series with the radiation resistance R.sub.a, which is roughly the same magnitude as X. These higher impedance elements reduce the sensitivity of the transducer for medical ultrasound scanning. The higher impedance element creates an impedance mismatch with conventional transmit circuitry in transmit mode, thus reducing acoustic power transmitted into the patient's body. In receive mode, the high impedance array element suffers significant losses when trying to drive conventional coaxial cable characterized by its capacitance/length.
It has been a significant challenge for the ultrasound community to design and fabricate linear phased arrays for medical ultrasound over the past two decades. Three performance characteristics have established conventional size and geometry of the transducer array elements: (1) the elements have sufficient angular sensitivity to steer the phased array over a .+-.45.degree. sector angle; (2) The arrays suppress grating lobe artifacts by fine inter-element spacing; and (3) the width of each rectangular element is small compared to the transducer thickness to remove parasitic lateral mode vibrations from the desired transducer pass band. Adherence to these performance characteristics have produced linear arrays having long narrow elements which are sized to be less than one wavelength wide for the ultrasonic frequencies used in tissue imaging, (e.g., &lt;0.3 mm wide .times.10 mm long at 3.5 MHz).
Two dimensional (N.times.M) transducer arrays are believed to hold promise in improving clinical image quality in future diagnostic ultrasound equipment. An immediate clinical application of 2-D phased arrays is the reduction of image slice thickness by focusing in the elevation plane perpendicular to the scanning dimension. An additional application of 2-D transducer arrays is the correction of phase aberrations introduced across the transducer aperture by tissue inhomogeneities. These aberrations occur in two dimensions, so 2-D arrays combined with the proper phase correction signal processing can restore diagnostic image quality. In addition to improving conventional ultrasound B-scan image quality, two-dimensional transducer arrays should assist in the development of new modes of ultrasound imaging. Projected new techniques include: (1) presentation of simultaneous orthogonal B-mode scans; (2) acquisition of several B-scans electronically steered in the elevation direction; (3) development of high-speed C-scans; and (4) high-speed volumetric ultrasound scanning to enable real time three-dimensional imaging and volumetric, angle-independent flow imaging. With known technology, these techniques can only be implemented with 2-D array transducers.
Unfortunately, the design and fabrication problems of one-dimensional transducer arrays become almost overwhelming when extended to a two dimensional array, in which case the element size may be less than 0.2 mm .times.0.2 mm for more than 1000 elements in the array. There are two significant obstacles which limit the use of 2-D transducer arrays. First, a simple fabrication method for the electrical connections to such array elements, which can be less than one ultrasound wavelength on a side, is not known. Second, it is very difficult to achieve adequate sensitivity and bandwidth from such small elements.
In the last 15 years there have been several descriptions of prototype 2-D array transducers for medical ultrasonic imaging, but the resulting products were acoustically unsuitable for modern medical ultrasound imaging procedures.
Two dimensional arrays also have been confronted with the problem of high electrical impedance in transducer elements. Two-dimensional arrays have been developed in two geometries. A typical geometry for a 4.times.32 array transducer is designed for focusing (but not steering) in the elevation direction and for correction of phase aberrations in two dimensions. Such transducers have been called 1.5-D arrays. For a transducer array of this design, each element typically exhibits complex impedance, the magnitude of which is approximately 1000.OMEGA. at a resonance of 3.5 MHz; this complex impedance causes an electrical impedance mismatch and the accompanying sensitivity decrease which are more severe than seen in linear arrays. Elements in full 2-D arrays which can steer the ultrasound beam in azimuth as well as elevation may be smaller than 0.2 mm.times.0.2 mm; these elements exhibit a complex electrical impedance having a magnitude of approximately 5000.OMEGA. or greater, so sensitivity is further reduced. Thus, for 1.5-D and 2-D arrays, the development of suitable piezoelectric materials are critical to improved sensitivity.
Unfortunately, the piezoelectric ceramics as described in the prior art are ill suited for such transducers. 1.5-D and 2-D arrays are commonly fabricated by dicing a single piezoelectric chip in two directions with a kerf width as small as 0.01 mm.
In an attempt to address the problem of high electrical impedance in linear arrays, U.S. Pat. No. 4,958,327 to Saitoh et al., teaches the concept of a multi-layer ceramic piezoelectric material consisting of K layers laminated in parallel electrically but in series acoustically. For K layers of uniform thickness, the capacitance of each element is increased by K.sup.2 ; this capacitive increase reduces the electrical impedance of the element by K.sup.2 significantly improving transmit efficiency and receive mode sensitivity. However, the teaching of Saitoh is inapplicable to two dimensional arrays; as the electrode layers are short circuited on a side surface of an element, the concept is limited to elements with electrode layers having a surface on the periphery of the transducer, and cannot be used for the elements of the inner rows of two dimensional arrays.
Transducers may be developed using a piezoelectric substrate fabricated from a composite of a piezoelectric ceramic phase such as PZT and an inert phase such as a polymer epoxy. However, piezoelectric composites produce a lower relative dielectric constant than PZT alone due to the presence of the epoxy phase. The lower dielectric constant results in a lower capacitance and thus a higher electrical impedance than PZT. The higher electrical impedance has limited the use of PZT/epoxy composites in the small elements of steered linear phased arrays and two dimensional arrays. Moreover, the lower dielectric constant exacerbates the high impedance problem of two dimensional array elements described above.
In view of the foregoing, it is an object of the present invention to provide a transducer array having decreased acoustic impedance, decreased electrical impedance and suppressed lateral mode.
An additional object of the present invention is to provide a method of making a transducer chip having decreased acoustic impedance, decreased electrical impedance and suppressed lateral mode.
It is a further object of the invention to provide an ultrasound transducer array which contains such a transducer chip.
It is an additional object of the invention to provide ultrasound diagnostic devices which utilize a transducer array as described.